Enzyme-based biosensors are devices in which an analyte-concentration-dependent biochemical reaction signal is converted into a measurable physical signal, such as an optical or electrical signal. Such biosensors are widely used in the detection of analytes in clinical, environmental, agricultural and biotechnological applications. Analytes that can be measured in clinical assays of fluids of the human body include, for example, glucose, lactate, cholesterol, bilirubin and amino acids. The detection of analytes in biological fluids, such as blood, is important in the diagnosis and the monitoring of many diseases.
Biosensors that detect analytes via electrical signals, such as current (amperometric biosensors) or charge (coulometric biosensors), are of special interest because electron transfer is involved in the biochemical reactions of many important bioanalytes. For example, the reaction of glucose with glucose oxidase involves electron transfer from glucose to the enzyme to produce gluconolactone and reduced enzyme. In an example of an amperometric glucose biosensor, glucose is oxidized by oxygen in the body fluid via a glucose oxidase-catalyzed reaction that generates gluconolactone and hydrogen peroxide, whereupon the hydrogen peroxide is electrooxidized and correlated to the concentration of glucose in the body fluid. (Thome-Duret, V., et al, Anal. Chem. 68, 3822 (1996); and U.S. Pat. No. 5,882,494 of Van Antwerp.) In another example of an amperometric glucose biosensor, the electrooxidation of glucose to gluconolactone is mediated by a polymeric redox mediator that electrically “wires” the reaction center of the enzyme to an electrode. (Csoregi, E., et al., Anal. Chem. 66, 3131 (1994); Csoregi, E., et al., Anal. Chem. 67, 1240 (1995); Schmidtke, D. W., et al., Anal. Chem. 68, 2845 (1996); Schmidtke, D. W., et al, Anal. Chem. 70, 2149 (1998); and Schmidtke, D. W., et al, Proc. Natl. Acad. Sci. U.S.A. 95,294 (1998).)
Amperometric biosensors typically employ two or three electrodes, including at least one measuring or working electrode and one reference electrode. In two-electrode systems, the reference electrode also serves as a counter-electrode. In three-electrode systems, the third electrode is a counter-electrode. The measuring or working electrode is composed of a non-corroding carbon or a metal conductor and is connected to the reference electrode via a circuit, such as a potentiostat.
Some biosensors are designed for implantation in a living animal body, such as a mammalian or a human body, merely by way of example. In an implantable amperometric biosensor, the working electrode is typically constructed of a sensing layer, which is in direct contact with the conductive material of the electrode, and a diffusion-limiting membrane layer on top of the sensing layer. The sensing layer typically consists of an enzyme, an enzyme stabilizer such as bovine serum albumin (BSA), and a crosslinker that crosslinks the sensing layer components. Alternatively, the sensing layer consists of an enzyme, a polymeric mediator, and a crosslinker that crosslinks the sensing layer components, as in the above-mentioned “wired-enzyme” biosensor.
In an implantable amperometric glucose sensor, the membrane is often beneficial or necessary for regulating or limiting the flux of glucose to the sensing layer. By way of explanation, in a glucose sensor without a membrane, the flux of glucose to the sensing layer increases linearly with the concentration of glucose. When all of the glucose arriving at the sensing layer is consumed, the measured output signal is linearly proportional to the flux of glucose and thus to the concentration of glucose. However, when the glucose consumption is limited by the kinetics of chemical or electrochemical activities in the sensing layer, the measured output signal is no longer controlled by the flux of glucose and is no longer linearly proportional to the flux or concentration of glucose. In this case, only a fraction of the glucose arriving at the sensing layer is consumed before the sensor becomes saturated, whereupon the measured signal stops increasing, or increases only slightly, with the concentration of glucose. In a glucose sensor equipped with a diffusion-limiting membrane, on the other hand, the membrane reduces the flux of glucose to the sensing layer such that the sensor does not become saturated and can therefor operate effectively within a much wider range of glucose concentration.
More particularly, in these membrane-equipped glucose sensors, the glucose consumption rate is controlled by the diffusion or flux of glucose through the membrane rather than by the kinetics of the sensing layer. The flux of glucose through the membrane is defined by the permeability of the membrane to glucose, which is usually constant, and by the concentration of glucose in the solution or biofluid being monitored. When all of the glucose arriving at the sensing layer is consumed, the flux of glucose through the membrane to the sensing layer varies linearly with the concentration of glucose in the solution, and determines the measured conversion rate or signal output such that it is also linearly proportional to the concentration of glucose concentration in the solution. Although not necessary, a linear relationship between the output signal and the concentration of glucose in the solution is ideal for the calibration of an implantable sensor.
Implantable amperometric glucose sensors based on the electrooxidation of hydrogen peroxide, as described above, require excess oxygen reactant to ensure that the sensor output is only controlled by the concentration of glucose in the body fluid or tissue being monitored. That is, the sensor is designed to be unaffected by the oxygen typically present in body fluid or tissue. In body tissue in which the glucose sensor is typically implanted, the concentration of oxygen can be very low, such as from about 0.02 mM to about 0.2 mM, while the concentration of glucose can be as high as about 30 mM or more. Without a glucose-diffusion-limiting membrane, the sensor would become saturated very quickly at very low glucose concentrations. The sensor thus benefits from having a sufficiently oxygen-permeable membrane that restricts glucose flux to the sensing layer, such that the so-called “oxygen-deficiency problem,” a condition in which there is insufficient oxygen for adequate sensing to take place, is minimized or eliminated.
In implantable amperometric glucose sensors that employ wired-enzyme electrodes, as described above, there is no oxygen-deficiency problem because oxygen is not a necessary reactant. Nonetheless, these sensors require glucose-diirusion-limiting membranes because typically, for glucose sensors that lack such membranes, the current output reaches a maximum level around or below a glucose concentration of 10 mM, which is well below 30 mM, the high end of clinically relevant glucose concentration.
A diffusion-limiting membrane is also of benefit in a biosensor that employs a wired-enzyme electrode, as the membrane significantly reduces chemical and biochemical reactivity in the sensing layer and thus reduces the production of radical species that can damage the enzyme. The diffusion-limiting membrane may also act as a mechanical protector that prevents the sensor components from leaching out of the sensor layer and reduces motion-associated noise.
There have been various attempts to develop a glucose-diffusion-limiting membrane that is mechanically strong, biocompatible, and easily manufactured. For example, a laminated microporous membrane with mechanical holes has been described (U.S. Pat. No. 4,759,828 of Young et al.) and membranes formed from polyurethane are also known (Shaw, G. W., et al., Biosensors and Bioelectronics 6, 401 (1991); Bindra, D. S., et al., Anal. Chem. 63, 1692 (1991); Shichiri, M., et al., Horm. Metab. Res., Suppl. Ser. 20, 17 (1988)). Supposedly, glucose diffuses through the mechanical holes or cracks in these various membranes. Further by way of example, a heterogeneous membrane ‘with discrete hydrophobic and hydrophilic regions (U.S. Pat. No. 4,484,987 of Gough) and homogenous membranes with both hydrophobic and hydrophilic functionalities (U.S. Pat. Nos. 5,284,140 and 5,322,063 of Alien et al.) have been described. However, all of these known membranes are difficult to manufacture and have inadequate physical properties.
An improved membrane formed from a complex mixture of a diisocyanate, a diol, a diamine and a silicone polymer has been described in U.S. Pat. Nos. 5,777,060 (Van Antwerp), 5,786,439 (Van Antwerp et al.) and 5,882,494 (Van Antwerp). As described therein, the membrane material is simultaneously polymerized and crosslinked in a flask; the resulting polymeric material is dissolved in a strong organic solvent, such as tetrahydroforan (THF); and the resulting solution is applied onto the sensing layer to form the membrane. Unfortunately, a very strong organic solvent, such as THF, can denature the enzyme in the sensing layer and also dissolve conductive ink materials as well as any plastic materials that may be part of the sensor. Further, since the polymerization and crosslinking reactions are completed in the reaction flask, no further bond-making reactions occur when the solution is applied to the sensing layer to form the membrane. As a result, the adhesion between the membrane layer and sensing layer may not be adequate.
In the published Patent Cooperation Treaty (PCX) Application bearing International Publication No. WO 01/57241 A2, Kelly and Schiffer describe a method for making a glucose-diffusion-limiting membrane by photolytically polymerizing small hydrophilic monomers. The sensitivities of the glucose sensors employing such membranes are widely scattered, however, indicating a lack of control in the membrane-making process. Further, as the polymerization involves very small molecules, it is quite possible that small, soluble molecules remain after polymerization, which may leach out of the sensor. Thus, glucose sensors employing such glucose-diffusion-limiting membranes may not be suitable for implantation in a living body.